This invention relates to x-ray detectors and more particularly to x-ray detectors which are useful in computerized tomography applications.
Computerized tomography scanners are medical diagnostic instruments in which the subject is exposed to a relatively planar beam or beams of x-ray radiation whose intensity varies in a direct relationship to the energy absorption along a plurality of subject body paths. By measuring the x-ray intensity, that is, the x-ray absorption, along these paths from a plurality of different angles, that is, different views, an x-ray absorption coefficient can be computed for various areas in the plane of the body through which the radiation passes. These areas typically correspond to an approximately square portion having dimensions of approximately 1 mm.times.1 mm. These absorption coefficients are used to form a display of the bodily organs intersected by x-ray beams. An integral and important portion of these scanners is the x-ray detector which receives the x-ray radiation which has been modulated by passage through the particular body under study. The detector functions to convert the resultant x-ray intensity level information to electrical signals which are then typically converted to a digital form to be processed by digital computer means which generates the absorption coefficients in a form suitable for display on a cathode ray tube screen or other, permanent media.
Conventional detectors used in computerized tomography are of two basic kinds. The first kind employs a scintillator crystal material while the second employs high pressure xenon gas as the detection medium.
In the scintillator crystal detector, a plurality of scintillator crystal slabs are separated by collimating plates comprising material such as tungsten or tantalum. The detectors based on crystal scintillators exhibit certain desirable properties, namely, they have a high overall quantum detection efficiency (better then 90 percent) and also exhibit a low attenuation length. For example, approximately 30 percent of the impinging x-ray photons from a 73 KeV x-ray beam are absorbed in a length of approximately 1 mm. In other words, approximately 97 percent of the impinging x-ray photons are absorbed in three "stopping distances", that is, approximately 3 mm. However, crystal scintillators exhibit other properties which are undesirable in computerized tomography applications. For example, crystal scintillators generally exhibit a poor spectral linearity. This is particularly important in the case of x-ray detectors since x-ray beams cannot, at present, be produced which exhibit a monochromatic frequency. Additionally, certain crystal scintillator materials such as cesium iodide (CsI) exhibit unsatisfactory afterglow characteristics. That is to say, certain crystal scintillator materials still produce optical wavelength radiation 100 microseconds after excitation. These afterglow characteristics render certain of these crystal scintillator materials unsatisfactory for rapid scanning which is particularly desirable in the event that moving bodily organs such as the heart and lungs and even the gastrointestinal tract are being studied. Moreover, solid crystal scintillator materials exhibit certain other problems. Among these are potential nonuniformities from crystal imperfections or machining scratches or cracks; quantum detection efficiency losses from gaps between crystals and adjacent collimator plates; and detector failure if just 1 of up to 1,000 crystals becomes dislodged or mispositioned; and view-to-view crosstalk from the fluorescent decay tails, especially for continuous, as opposed to pulsed, x-ray beam operation.
The other commonly used kind of x-ray detector for computerized tomography applications is the xenon detector. In this detector, closely spaced electrically conductive collimating plates are disposed in a housing containing xenon gas at a pressure of approximately 25 atmospheres. Adjacent plates are maintained at opposite high electric potentials and the operation of the detector depends upon the production of electrons from the xenon gas by the impinging x-ray photons. These electrons drift under the influence of the applied voltage to the anode detector plate and a current results depending upon the intensity of the impinging x-ray beam. These xenon detectors respond to impinging x-ray radiation in several microseconds and exhibit an overall quantum detection efficiency of less than approximately 70 percent. Moreover, they exhibit a 30 percent stopping distance of approximately 16 mm, which is longer than desired, but not prohibitively long. Xenon detectors exhibit excellent spatial homogeneity since xenon is a gaseous medium but like crystal scintillators, exhibit poor spectral linearity. For example, the ratio of the attenuation length for 110 KeV photons to the attenuation length for 40 KeV photons is approximately 15:1, as is also the case for crystal scintillators, such as CsI and BGO (bismuth germanate, Bi.sub.4 Ge.sub.3 O.sub.12). There also other problems associated with the high pressure xenon detector. In particular, because of the need for a high gas pressure, a housing is required with a front wall comprising a material which is relatively transparent to x-ray radiation but which is structurally sound. The housings typically comprise aluminum, which while being structurally sound, does absorb a certain number of x-ray photons which are not therefore subsequently used to indicate the x-ray intensity level. Additionally, minute movements of the collimator plates which result from movement of the scanner gantry or from other vibrational sources, often produce microphonic noise which can produce image artifacts. This microphonic noise arises at least in part because of the necessity of maintaining adjacent plates at opposite but high electric potentials.
As used above, the term "quantum detection efficiency" refers to the fractional number of x-ray photons which are absorbed photoelectrically in the detection medium, whether the medium be a crystal or a gas. Another term of merit which is often applied to these x-ray detectors is the "conversion efficiency" which includes losses due to the photodetector system that is employed. For example, the conversion efficiency for a gridded xenon detector system is approximately 8 percent while the conversion efficiency for a scintillator crystal employing CsI is approximately 15 percent.